MRI Compatible 3-D Intracardiac Echography Catheter and System

ABSTRACT

An intracardiac imaging system has an intracardiac echography catheter with an internal volume, a proximal end and a distal end. The catheter includes an atraumatic tip disposed on the distal end of the catheter, a pair of inductively coupled coils proximal the atraumatic tip, at least one CMUT on CMOS volumetric imaging chip disposed between the pair of coils, and a cable lumen disposed within the volume and configured to small number of electrical connections due to significant multiplexing in the CMUT on CMOS chip. The catheter can be made of MRI compatible materials and can include active cooling channels. The CMUT on CMOS chip has a plurality of Tx elements transmitting imaging pulses, a plurality of Rx elements, disposed on the chip to have a large aperture and a plurality of electronics interfacing with the Tx elements for beamforming and the Rx elements to produce radio frequency output signals.

CROSS-REFERENCE TO RELATED APPLICATION

This application claims priority to U.S. Provisional Application Ser.No. 61/882,371 filed Sep. 25, 2013. The entirety of the application isincorporated herein by reference.

Government License Rights

This invention was made with government support with Grant No. EB010070that was awarded by National Institutes of Health (NIH). The governmenthas certain rights in the invention.

FIELD OF THE INVENTION

The invention related to a system for intracardiac imaging with anintracardiac echography catheter utilizing CMUT on CMOS technology forvolumetric ultrasound imaging. The catheter can be operated in an MRIsystem as well as in an X-Ray system to guide intracardiac interventionsin real time. The CMUT on CMOS technology is used to integrate transmit(Tx) electronics into the catheter tip and heavily multiplexing thereceive (Rx) elements. This results in a catheter with small number ofcables significantly reducing the heating of the catheter under large RFsignals used for MRI. Reduction of number of cables also reduces thecross sectional area required for electrical connections and makes roomfor active cooling of the catheter. The catheter also has integratedmarkers for tracking its position under MRI.

BACKGROUND

Symptomatic adult and pediatric structural heart disease (such asvalvular heart disease or cardiac septal defects) affects more than 2.9%of the US population, not including cardiomyopathies and rhythmdisorders. Because of procedural morbidity, only a minority are selectedfor surgical relief of symptoms. Nonsurgical repair of structural heartdisease is possible using image guidance and newer devices such astranscatheter aortic valves, mitral valve repairs, and intracardiacoccluders. Most are guided by X-ray fluoroscopy and adjunctive 2Dintracardiac or 3D transesophageal echocardiography (“TEE”). Whileavailable transesophageal and intracardiac echo systems are suitable toassess target pathology immediately before and after treatment, they areunsuitable to guide catheter manipulations during therapeuticprocedures. Catheters and target pathology constantly move outside the2D slices and limited 3D volumes depicted by current echo systems, whichalso are constrained by interposed lung and bone or by esophageal accessroute. As a result, operators are forced to use X-ray fluoroscopy toguide catheter manipulation in contemporary repair of complex atrial andventricular septal defects, valve leaflets, valve replacement,paravalvular leak, and left atrial appendage closure; operators muststruggle visually to integrate 2D images into a mental image of anatomiccontext during key steps of protracted and occasionally unsuccessfulprocedures. Moreover, current 3D TEE probes, although shown to be usefulin repair of septal defects, are not small enough for young children.Miniaturization of ultrasound probes to provide uninterrupted real-timefull-volume intraprocedural three-dimensional en face depiction ofcardiac pathology and catheter devices would represent a dramaticadvance in image-guided intervention

SUMMARY

The goal of this invention is to dramatically enhance image guidance ofcomplex catheter-based cardiovascular treatments, to avoid radiationexposure especially in children, to allow current procedures to beperformed more safely and efficiently, and to enable novel proceduresthat otherwise might require surgical repair. Common procedures such asatrial septal defect closure, and emerging procedures such as closure ofventricular septal defects and paravalvular leak, future repair of valveleaflets, transcatheter valve replacement, and emerging left atrialappendage closures can be difficult, protracted, or unsuccessful becauseof limitations of available interventional catheter devices but alsobecause of inadequate image guidance.

Commercially available 2D and limited-volume 3D intracardiac ultrasoundcatheters do not provide suitable full-volume en face images to depictcomplex cardiac structures in real time, do not adequately depictreal-time navigation of catheter tips and shafts, and require adjunctiveX-ray guidance. Several 3D catheters with 2D arrays under developmentuse over 200 electrical connections limiting size and flexibility, andprohibiting operation under MRI (magnetic resonance imaging). Disclosedbelow is the capability to build an ultra-miniature ultrasoundsystem-on-a-chip that provides realtime full-volume 3D ultrasound withvery few external electrical connections. This can be implemented as alowprofile steerable intracardiac catheter and that further can beimplemented by design for operation under either MRI or X-ray.

To reach this goal, one advance is MRI catheterization as aradiation-free alternative to X-ray. However, this trades the safety oflower radiation emissions at the expense of real-time spatialresolution. An intracardiac echography (“ICE”) operation during MRI candramatically advance or even revolutionize the capabilities oftranscatheter therapy by enabling completely radiation-free non-surgicalcatheter navigation, depiction of anatomic context, device repair, novelprocedures, and assessment of success and complications, in children andadults.

Full volumetric ICE poses significant challenges even apart from MRIsafe operation. Ideally, a fully populated 2D matrix phased array with100 nm×100 nm or smaller elements should be used for 3D ICE. Traditionaldesigns require large numbers of transmission cables, which cause anumber of difficulties. Some of the problems are, prohibitivemanufacturing complexity and cost, prohibitive form factor forintracardiac catheters, and (incidentally) increased propensity to RF(radio frequency)-induced heating of metal conductors during an MRI.

The tight space constraints of ICE catheters can also precludeintegration of electronics with conventional 1D or 2D matrixpiezoelectric arrays needed to improve the signal-to-noise ratio (“SNR”)and to implement microbeamformer concepts and thereby enable 3D TEEprobes. This leaves motor driven 1D arrays or swept aperture techniquesas the only available alternative without increasing the cable count,and such systems suffer from inadequate view angles and large slicethickness in the elevation direction.

Piezoelectric micromachined ultrasonic transducer (“pMUT”) andcapacitive micromachined ultrasonic transducer (“CMUT”) technologiesprovide more robust fabrication methods for 2D matrix arrays as comparedwith traditional piezoelectrics and both have been shown to haveadequate performance for volumetric imaging with approximately 200elements and same number of cables. Ring annular array structuresfurther reduce the element count and can still provide 3D imageguidance, for example along with integrated RF ablation capability.However, such implementations have small active array areas,exacerbating the compromise between penetration depth and tolerance totissue motion which is critical in ICE.

In all these approaches, even when flip-chip technology and complexthrough-silicon electrical connections are used for CMUT ring array-CMOS(complementary metal-oxide-semiconductor) electronics integration. Eacharray element is still connected to the imaging system with a separatecable resulting in a catheter with more than 70 cables. Therefore,real-time 3D ICE implementation, which requires full volumetric datacollection from less than 10 array transmit firings due to fast tissuemotion and miniaturization-driven reduction in the number of datatransmission lines, requires a different level of system complexityimplemented at the catheter tip, even apart from the requirements of MRIsafety

Low temperature fabrication can be used to build CMUT arrays on the samesilicon substrate as the CMOS electronics. This approach, calledCMUT-on-CMOS, enables integration of full 3D transmit and low noisereceive frontend electronics as well as RF output multiplexing on asingle silicon chip to reduce the cable count. An example of thistechnology utilized in the present invention can achieve thermalmechanical noise limited detection and real-time 3D imaging at 20 MHzwith a 1.4 mm diameter 104 element ring array with only 13 electricalconnections.

As with other ring arrays, that particular system also presented atradeoff between motion artifacts and penetration depth. Penetrationdepth can be improved by utilizing a larger transmit array areaavailable for a side looking 3D ICE array and implementing on-chip codedexcitation schemes as discussed below. By massive on-chip multiplexingof high SNR receive signals over a few cables, image data acquisitiontime can be reduced a few firings to minimize motion artifacts.Therefore the CMUT-on-CMOS approach, along with innovative on-chipbeamforming and massive multiplexing, provides a unique platform forfull-volume real-time 3D ICE, and the only one suitable for MRI safeoperation.

The expected benefits of MRI plus ICE guided structural heartinterventional procedures are manifold. Enhanced visualization promisesto simplify and shorten current procedures to enhance success, reducecomplications, and reduce cost. Enhanced guidance combined with newerdevices can enable catheter alternatives to surgery such as non-surgicalextra-anatomic bypass (e.g., Glenn shunt, modified Blalock-Taussigshunt) to reduce the steps of Norwood palliation; simplified repair ofmultifenestrated muscular and of membranous ventricular septal defect byvirtue of en-face imaging during device manipulation; leaflet graspingprocedures for neochordal implantation to treat degenerative andfunctional mitral valve regurgitation; and leaflet or annular orsubvalvar plication or augmentation of the mitral and tricuspid valves.At present all of these procedures are challenging or unrealistic absentdirect surgical visualization. Supine 3D TEE for procedure guidanceusually requires prolonged and costly general anesthesia; intracardiac3D ICE can avert this need and thereby reduce staffing cost (by 1-2physicians) and risk. En-face imaging of an atrial septal defect (“ASD”)potentially may enhance sizing for device selection to avert rare butcatastrophic erosion after implantation of an Amplatzer Septal Occluder,and of dynamic sizing of the ostium of the left atrial appendage mayovercome the limitations of available alternatives including 3D TEE.

The 3D real-time full-volume MRICE catheter can, for the first time,allow routine ultrasound guidance of catheter manipulation duringprocedures rather than just inspecting the baseline pathology andresults of repair. It can allow universal real-time en-face depiction oftarget pathology without the constraints of bone and lung windows(transthoracic) and limited probe positioning (transesophageal) incurrent technology. For the first time it can enable completelyradiation-free catheter navigation and depiction of larger anatomiccontext and tissue characterization using real-time MRI instead of X-rayfor catheter navigation. Even without operation under MRI, real-timefull volume 3DICE with higher probe frequency would represent afundamental advance for conventional X-ray catheterization. It canenable new procedures not currently possible without surgery, such asnon-surgical mitral neochordal implantation and direct mitralannuloplasty and can greatly simplify complex structural heartinterventional procedures such as paravalvular leak repair,postinfarction and congenital muscular VSD repair, left atrial appendageclosure, atrial and ventricular myocardial ablation procedures forrhythm disorders.

BRIEF DESCRIPTION OF THE DRAWINGS

This invention is described with particularity in the appended claims.The above and further aspects of this invention may be better understoodby referring to the following description in conjunction with theaccompanying drawings, in which like numerals indicate like structuralelements and features in various figures. The drawings are notnecessarily to scale, emphasis instead being placed upon illustratingthe principles of the invention.

The drawing figures depict one or more implementations in accord withthe present teachings, by way of example only, not by way of limitation.In the figures, like reference numerals refer to the same or similarelements.

FIG. 1 is a distal-side perspective view of a three-dimensional MRICEcatheter;

FIG. 2 is a perspective view of several of the possible CMUT-on-CMOSarray examples for 3D ICE;

FIG. 3 is an example high level circuit schematic of the on-chipelectronics for massively parallel Rx RF data transfer and on-chip Txbeamforming;

FIGS. 4A and 4B are the handle and a cross-section of the proximal end,respectively, of an example of a three-dimensional MRICE catheter;

FIG. 5 is a block diagram of an example of a real-time imaging system;

FIG. 6 is an example high level circuit schematic of a portion of theon-chip electronics for another example for massively parallel RF datatransfer;

FIG. 7 is an example high level circuit schematic of a portion of theon-chip electronics for demodulating the combined signal of FIG. 6;

FIG. 8 is a further example high level circuit schematic of the on-chipelectronics for massively parallel Rx RF data transfer;

FIG. 9 illustrates a bock diagram of an example of analog OrthogonalFrequency Division Multiplexing (OFDM);

FIG. 10 illustrates a layout of an example of an 8 channel OFDM circuitfor ICE;

FIG. 11 is a schematic of a capacitive feedback TIA;

FIG. 12 is a schematic of the single to differential converter;

FIG. 13 is a schematic of a example of a biquad structure used fordesigning gm-C low pass and band-pass filters;

FIG. 14 is a schematic of an I/Q passive mixer;

FIG. 15 illustrates a simulated frequency response of the capacitivefeedback TIA;

FIGS. 16A and 16B illustrate the transient simulation results of thesingle to fully differential conversion circuit;

FIG. 17 illustrates a simulated Frequency response of 4th Order LPF;FIG. 18 illustrates a simulated frequency response of the designedband-pass filters;

FIGS. 19A-19C illustrate a transient simulation of analog OFDM with 2 7MHz Gaussian pulses; and

FIGS. 20A-20C illustrate a spectrum of 8 channel analog OFDM signals.

DETAILED DESCRIPTION

FIG. 1 illustrates a full-volume 3D MRICE (“magnetic resonanceintracardiac echography”) catheter 100 with a tip including at least oneCMUT-on-CMOS chip for Rx multiplexing and Tx beamforming 102. Thisexample can take advantage of the CMUT-on-CMOS technology for 2D arraydesign flexibility and on-chip transmit/receive beamforming electronicsintegration, recent advances in FPGA and graphical processing unit (GPU)based real-time ultrasound image processing, and MRI compatible catheterdesign and implementation. The 3D MRICE catheter 100 is MRI safe. The 3DMRICE catheter 100 can range in sizes from 6 French to about 10 French(French (Fr)=Diameter (mm)*3) and can have a 2-axis deflection point 104to provide access and views for most intracardiac operations. Anotherelement of the 3D MRICE catheter 100 is the inductively coupled coils106, to provide visibility and to navigate the catheter tip 110 in spaceunder MRI.

The 3D MRICE catheter 100 is similar in mechanical properties andsteerability to current 2D ICE, and thus less likely to fail clinically.One example, as illustrated in FIG. 2, the CMUT-on-CMOS volumetricimaging array 102 can have a large area (7 mm×2.8 mm) to provide anarray size with adequate spatial resolution, and large active area foracoustic power output for operation in 5-12 MHz range. This is alsocritical for electronics integration since spatially and temporallycoded transmit signal generation and low noise received signal detectionfunctionality can be contained in a single chip 102. One can also use aCMUT-on-CMOS in a multiple stacked chip configuration to have more areafor electronics under the same array area as described in as describedin U.S. Pat. No. 8,766,459, the disclosure of which is incorporatedherein by reference.

As an example, the imaging can be done over a 90°×90° field of view(“FOV”) at 5 cm, and narrowing to 45°×45° at 15 cm. This is one exampleof a desired spatial range of most ASD, ventricular septal defect(“VSD”), left atrial appendage (“LAA”) occlusion, and mitral procedures.Within this framework, 3D MRICE catheter 100 can provide: (1) Collectionof full volumetric ultrasound image data over 5-15 cm penetration depthwith less than 10 transmit firings; (2) high information rate over fewtransmission lines; (3) catheter and operational design to avoid MRIinterference, and (4) thermal management of RF catheter heating.

To provide these benefits, the key features of the invention are: 1.CMUT-on-CMOS technology implementing large aperture 2D receive arrayswith more than 100 elements and about 100 nm×100 nm element size forlarge FOV and integrated low noise electronics to obtain high SNR.

2. On-chip electronics that allow for massively parallel RF datatransfer (in an example, greater than 200 MHz bandwidth per line) tocapture volumetric image data in few transmit firings.

3. On-chip electronics and backend processing strategies for volumetricimaging with minimal motion artifacts, from simply defocused/focused,temporally coded defocused/focused to spatially coded multiplane phasedarray transmit beamforming implementation. This feature, along withmassively parallel RF data transfer, can allow up to 50× reduction incable count as compared to conventional cabling.

4. A 3D ICE catheter with as few as 14 transmission lines to minimize RFheating under MRI, and concurrent or (if necessary) coordinated MRI RFexcitation to minimize MRI interference.

5. A closed-loop actively cooled MRI safe ICE catheter design usingmaterials and techniques to minimize MRI artifacts and RF heating.

6. Inductively or conductively coupled marker coils for cathetertracking under MRI with minimized RF heating.

Sample array designs for 3D MRICE development are summarized in FIG. 2.In some examples, these are 2D CMUT-on-CMOS Arrays and On-chip CodedBeamforming that allow for 3D MRICE. These exemplary designs aim toobtain volumetric image data with minimum number of firings and achievethe required SNR for 5-15 cm imaging range. Initial designs can have90°×90° FOV at approximately 5 cm and 45°×45° FOV at approximately 15cm, selectable electronically. To maximize the active area, in all arraydesigns, the full silicon surface is covered by CMUT elements overlayingthe CMOS electronics. In this example, the 192 element receiving (“Rx”)arrays 200 are placed at the periphery to maximize the Rx aperture. Insome other embodiments the Rx array can have cross or plus shapes tocover the entire aperture with smaller number of elements. In general,well-known 2-D sparse array designs can be utilized to form the Rxarrays. Also, all of the Rx elements 200 read out in parallel during aone or two Tx firings, as described later. The example of ICE 1facilitates Rx electronics and real-time imaging system. A transmitting(“Tx”) array 202 can be kept simple in ICE 1 204. The Tx array 202 canbe driven by a short pulse or coded excitation to further improve SNR.In an example, a 14 dB gain in SNR with 13 bit, 2 cycle, 2.5 ns codescan be achieved with dual-ring CMUT-on-CMOS arrays. The Tx array 202fires a defocused imaging pulse received by the Rx arrays 200.

Another example ICE 2 206 design improves the lateral resolution beyondcurrent 2D ICE arrays by Tx beamforming in both directions and operatingat 10 MHz center frequency. It can also achieve ˜2 mm slice thickness inelevation at 5 cm. Another example ICE 3 208 design can add phased arraycapability with spatial and temporal coding. In this example, an imageover a 2D plane can be obtained during each Tx firing. Up to 90 planeimages, which can be displayed in multiplanar format, can be collectedto form the 3D volume. In this case the volumetric image is formed planeby plane where image for each plane is collected using one or twotransmit firings Improved resolution in ICE 2 206 and ICES 208 can berealized using improved SNR from design improvements, coding, andincreasing the imaging frequency. In this example, these approachesenable full volumetric imaging with minimum motion artifacts. AlthoughDoppler flow is not considered, flow measurement over 2D planes can beimplemented as part of the real time imaging system using correlationtechniques over frames obtained during consecutive firings.

In other examples of the invention, the on-chip electronics with massiveRF multiplexing for fast full volume imaging overcomes one of thechallenges for on-chip electronics. The examples of the invention canovercome the difficulty of the parallel readout of 192 Rx channels over8 RF transmission lines during each firing. Overcoming this existinglimitation can reduce the total number of transmission lines. This canbe achieved by frequency division multiplexing (“FDM”) or time divisionmultiplexing (“TDM”) using interleaved samples from different Rxchannels on the same line. This same technique can be used to reduce thenumber of cables for 1-D ICE arrays for other purposes including makingthem suitable for use under MRI.

A multiplexing component 301 is illustrated by example using a FDMsolution can be analog Quadrature Amplitude Modulation (“QAM”), which isa form of frequency division multiplexing. An example of a FDM solutionusing QAM 300 is illustrated in the overall electronics schematic ofFIG. 3. At the output of each Rx element 302 there can be a metal oxidesemiconductor (“MOS”) feedback transimpedance amplifier (“TIA”) 304. TheMOS feedback TIA 304 can provide low noise, high bandwidth andadjustable gain. The gain of the TIAs 304 can be dynamically changed fortime gain compensation. The power consumption of each TIA in thisexample can be about 0.825 mW and the entire TIA array 306 consumes 132mW. Signals from the TIAs 304 can be fed to single balanced MOS Gilbertmixer 308 to shift their frequencies to implement the analog QAM. Thisexample can shift the TIA outputs so that with each carrier frequencyand its quadrature component can shift two TIA outputs. Thus, using 12high frequency carriers (e.g., 40, 80, 100, 120, 140, 160, 180, 200,220, 240, 260 and 280 MHz) and their quadrature component, 24 TIAoutputs can be shifted and sent out via single cable 310. These RFcables 310 can be chosen to have sufficient bandwidth and minimizedelectrical crosstalk. To avoid interference from MRI signals, thefrequencies around the MRI Larmor frequency (around 63.9 MHz for a 1.5Tsystem) may be avoided. The carrier signal can be generated by an onchip arbitrary waveform generator 311 or by frequency multipliers usingan external clock. After the mixer 308, each signal can be passedthrough a bandpass filter (“BPF”) 312 to suppress the side band andharmonics. The mixers 308 and BPF bank 312 consumes approximately 120 mWand 225 mW power, respectively, in this example. After the BPF stage312, signals can be added using analog adders 314 and sent off chipthrough analog buffers 316. The analog buffers 316 can consume about 70mW power, for example. A digital control circuit 318 can be used topower down of the receiver circuits while the catheter is not collectingdata. Therefore, in this example, the chip 300 can consume about 0.6Wpeak power on the receiver side.

The average power can be much lower since even in the ICE 2 206 design,the chip can be active for only 18 ms of a 50 ms duty cycle at 20frames/second. Even when the average power consumption of the Tx side isadded, the overall figure can be significantly lower than 3-4W consumedin 2D ICE catheters, again due to lower duty cycle. An on chiptemperature 320 sensor can be implemented for continuous monitoring ofMRI induced heating, and the chip can have a shut off feature when thetemperature exceeds 43° C.

On the transmit side, a beamforming component 350 can be used and isillustrated in an example having each CMUT transmitter element 352connected with a high voltage on chip pulser 354. To change the FOVdepth from 5 cm to 15 cm, the pulse repetition rate can be changed. Inone example, this can be done by using an on chip counter 354. Atemporal and spatial coded excitation sequence can be used in the ICEchips 300, the code can be stored on-chip using a flash memory array,floating gate arrays or can be generated using digital logic circuitsand a clock signal. For programming the on-chip flash memory, a fewextra cables can be required which can be cut off once the chip isprogrammed before mounting on the catheter 100.

In an example of ICE 1 204, DC voltages can be applied directly, or DCvoltages can be generated on chip from an AC input signal to improveelectrical safety and to further reduce the transmission line count to14. This represents 15×, 25× and 50× reduction in transmission linenumber as compared to traditional implementations of the ICE 1, ICE 2,and ICE 3 arrays 204, 206, 208, respectively, considering that ICE 3 308array has 704 elements.

FIGS. 4A and 4B illustrate an example of a 3D ICE catheter design forMRI safe operation. The sterile, single-use 3D MRICE catheter 400 shownin FIGS. 4A and 4B can have a working length of approximately 110 cm. Inan example, a 10 Fr catheter shaft 402 can have several biocompatiblepolymer layers 404 with different durometers. The metal content (i.e.pulling wires for tip deflection and braiding) can be minimized toreduce the RF induced heating risk during MRI, and to reduce artifactsusceptibility which may obscure nearby anatomy.

In an example, the proximal shaft 402 can be reinforced withnon-metallic fibers (e.g., Vectran® or Kevlar®) instead of Nitinol orMP35N alloy braiding wires. The non-metallic fibers can preservecatheter pushability and torquability. The catheter 300 can have acylindrical enclosure 108 to house the side looking CMUT-on-CMOS chip102, 204, 206, 208, and a round atraumatic distal tip 110. A semi-rigidpolymer or MRI compatible metals can be used in the enclosure bodydepending on the final design.

In a manufacturing example, the enclosure 108 can be designed using 3DCAD software (e.g., Wildfire 4.0, Pro-Engineer). A metal model from theCADs can be manufactured from thin wall nitinol tube using 4-axis lasermetal processing equipment (e.g., ProLas, Lasag Laser Industries) andthe polymer model can be formed using a rapid prototyping system(Uprint, Strasys Inc., MN).

Non-planar inductively coupled marker coils or loop coils 106 (that canprovide separate receive channels connected to a scanner via coaxialtransmission lines) can be embedded into both ends of the enclosuregroove 108, in order to impart unambiguous real-time MRI visibility andtrackability to the catheter 100. While the SNR of inductively coupledcoils can be orientation dependent, this example allows furtherminiaturization by eliminating coaxial transmission lines, which occupyvaluable space within the catheter shaft and which also would contributeto RF induced heating.

The example of the multi-lumen thermoplastic catheter shaft 402 shown inFIG. 4B can have lumen space for pulling wires for two-planedeflections. One set of wires can pass through anterior and posteriorcable lumens 406 and left-right cable lumens 408. The shaft can furtherhave lumens 410 for an ultrasound system cable, and open and closed loopsaline cooling lumens 412. A polytetrafluoroethylene (PTFE) liner can beused in the pull cable lumens 406, 408 to reduce friction during use.Different durometers of otherwise matching design can be used indifferent sections of the catheter to accommodate deflection. Anon-metallic braiding can be performed over the first shaft layer usinga vertical 16-head braiding machine (e.g., a K16 Vertical braider,Steeger). The braiding angle can change between 20 and 60 degrees (whichaffects the torquability) and braiding density can change between 35 and80 picks per inch (which affects the shaft stiffness and pushability)can be optimized in non-metallic fibers to resemble metallic braidedshaft mechanical properties. The final PEBAX (polyether block amide)layer can be applied using reflowing technique (e.g., 810 Shrink cycler,Beahm Designs) to create a smooth catheter surface.

A distal tip deflector mechanism 414 (see FIG. 4A) for two differentplanes can be designed and manufactured from vertebrated nitinol superelastic alloy. The deflection can occur for the distal 70 mm to achievethe desired 30-35 mm radius of curvature. The deflection angle can be upto 180 degrees in each direction. The non-metallic sets of pull-cablescan be fixed on the metal deflection mechanism and each of the pullingwire can be advanced into the dedicated lumen 406, 408 within thecatheter shaft. The CMUT-on-CMOS imaging chip 102 can be connected tothe signal carrying cables directly via wire bonding or solderingtechnique. The cables can be bundled to reduce the occupied volume andcan be advanced into the catheter shaft using dedicated micro lumens.The cables can be soldered to the custom design male connector at theproximal end of the catheter. There can be two dedicated lumens 410 forthe ultrasonic imaging array cables. The cable lumens 410 can besurrounded by cooling lumens 412 which can be used for closed-loopcirculating-liquid cooling. In some cases, an open loop cooling systemwhere the cooling fluid is disposed into the blood stream can be used.

The system can handle any RF induced heating of the transmission lineswhen under MRI. Although the CMUT-on-CMOS silicon chip 102 is notexpected to heat under MRI, the cooling can also remove the heatconducted to the chip through the solder connections. The two separatelumens 412 can converge at the distal end, and can connect to a rotarycirculation pump. The liquid circulation speed can be adjusted based onthe real time temperature measurement through embedded thermistor probelocated on the CMUT-on-CMOS chip 102, 204, 206, 208 in the 3D MRICEcatheter 300. The temperature data can be transmitted on one of the RFoutput cables when no imaging data is being collected and before thepower is turned off for the next frame. The 3D MRICE system can providethat data to the display located in the MR control room and also to thecooling system controller that adjusts the rotary circulation pump speedwithin predetermined range. Both the controller and the rotarycirculation pump can be located in the MRI control room. The temperaturedata can be projected to the MRI room for the operator's review. The 3DMRICE catheter handle 416 can provide dedicated buttons to controldeflection amount and direction for each plane with single-handedoperation. The proximal end of the handle also has dedicated ports to beconnected to the imaging equipment and the cooling pump.

The imaging system can include a graphics processing unit (“GPU”) basedreal-time 3D MRICE volumetric imaging and graphical user interface(“GUI”) for flexibility in implementing different beamforming schemesand image processing, as illustrated in FIG. 5. The system 500 can havean Electrocardiography (“ECG”) input 502 for synchronization anddisplay. The system 500 can also be synchronized with an MRI system forsimultaneous operation or , time multiplexed (interleaved) operation, ifneeded. Timing information can be transferred to the imaging chipthrough the clock and reset inputs 504. The temperature sensor 520 onthe CMUT-on-CMOS chip 522 can be readout over the RF lines once perframe with proper timing in order not to interfere with the RFultrasound signals.

The other relevant system and processing requirements of the ICE arrays204, 206, 208 are given in FIG. 2. A part of the system 500 can beimplemented through minor modifications to many generic ultrasoundimaging platforms available from commercial vendors like Verasonics orsome research platforms such as Ultrasound Array Research Platform(“UARP”) system developed in University of Leeds. The approaches toimplement the system 500 would use combinations of a field programmablegate array (“FPGA”) and a GPU to perform the digital computationsefficiently. The volumetric image rendering can be performed on a GPUusing application specific software which can utilize public resourcessuch as the Gadgetron Open Source software.

For real time volume rendering and multi-plane image reconstruction, anOpen Source framework for medical image reconstruction, the Gadgetron,which has recently been developed at the NHLBI and at Aarhus University,Denmark, can be utilized. Several previous projects have alreadydemonstrated that it is indeed possible to obtain the desired volumetricrendering rates on the GPU, and that high performance open sourcesoftware tools are available, as well as several tutorials from leadinggraphics conferences. The GUI for this application can resemblecommercial 3D TEE and can depict multiplanar 2D images andsurface-rendered 3D volumes. It also can allow 3D point-marker placementfor complex geometry assessment.

In a further example of the multiplexing component, the reduction ofcables of ICE can also be achieved by implementing on-chip highfrequency TDM. FIG. 6 illustrates TDM multiplexing component 303 forthis example. Here, the TIA 304 amplifies the CMUT current signals asbefore. However, the TIA's 304 output now can be dynamically changedwith a time-gain compensation (“TGC”) circuit 330 to account for timegain. Each of TGC's 330 output is sampled at more than the Nyquist rate(which is twice the bandwidth of the bandlimited channel) in asynchronized way with a time division multiplexing (“TDM”) switch 332.Further, the output can be sent via single path and the signal is sentout via output buffer 316. FIG. 7 illustrates one example of how acombined signal 360 can be recovered. The combined signal 360 can sentto a de-multiplexer (switch) 362 and then transmitted through a low passfilter 364.

In an example of ICE 1 204, DC voltages can be applied directly, or DCvoltages can be generated on chip from an AC input signal to improveelectrical safety and to further reduce the transmission line count to14. This represents 15×, 25× and 50' reduction in transmission linenumber as compared to traditional implementations of the ICE 1, ICE 2,and ICE 3 arrays 204, 206, 208, respectively, considering that ICE 3 308array has 704 elements.

FIGS. 8-21 illustrate an example of a multiplexing component 800 usingorthogonal frequency division modulation or multiplexing (“OFDM”), andspecifically analogue OFDM. In this method, in the multiplexing step,the message signals from consequent channels are mixed with sine andcosine signals at a carrier frequency and added, with the final signalexpressed as:

${X(\omega)} = {{\sum\limits_{{n = 0},1,\ldots,k}{{m_{{2n} + 1}(\omega)}*{\cos \left( \omega_{n} \right)}}} + {m_{{2n} + 2}*{\sin \left( \omega_{n} \right)}}}$

For demodulation, the received signal is mixed with orthogonal sine andcosine signals at the modulation frequency and then low pass filtered toget back the message signals as:

${{X(\omega)}*{\cos \left( \omega_{n} \right)}}\overset{LPF}{\rightarrow}m_{{2n} + 1}$${{X(\omega)}*{\sin \left( \omega_{n} \right)}}\overset{LPF}{\rightarrow}m_{{2n} + 2}$

FIGS. 8, 9 and 10 illustrate different descriptions of the OFDMmultiplexing component 800. FIG. 10 illustrates an integrated circuitcustom designed to multiplex the output of 8 CMUT array elements withcenter frequency 7 MHz and 80% fractional bandwidth for ICE application.Parallel readout of 8 CMUT signals over single RF transmission lineduring each firing was achieved by implementing on chip analog OFDM withmodulation frequencies of 40 MHz, 80 MHz, 120 MHz and 160 MHz. Thesefrequencies are chosen to provide enough separation between channels aswell as avoiding the 64 MHz MRI signal in a 1.5T system, as noted above.The IC was designed in 0.35 μm 4M2P TSMC process with supply voltage of3.3V and the layout of the circuit consumes 0.7×1.76 mm² area.

FIGS. 8 and 10 illustrate the sequence of components. The Rx elements802 transmit their output to a capacitive-feedback TIA 804 and thesingle ended signal of the TIA 804 can be converted to fullydifferential with a single to differential converter (“STD”) 806. A lowpass filter (“LPF”) 808 filters the signal outputted from the TIA 804and the STD 806 and sends it to a passive mixer 810 to modulate thesignal to an orthogonal carrier frequency. To eliminate the harmonicgenerated by the passive mixer 810, the signal is then put through aband pass filter 812 and then on to a buffer 814 before transmitting theoutput signal off the chip 800.

Specifics regarding some of the components of the OFDM multiplexingcomponent 800 are illustrated in FIGS. 11-14 and discussed below. FIG.11 illustrates the capacitive-Feedback TIA 804. It can be the firstcomponent of the receiver chain acting as a low-noise preamplifier.Common to all sensor front-end circuits, the first amplifier's noiseperformance in the receiver path determines the SNR of the entire signalpath. For its low input referred noise, high gain and high bandwidthfeature the capacitive feedback TIA 804 can be used to amplify theCMUT's output current. In this example, the TIA' s 804 simulatedbandwidth can be ˜20 MHz and Gain 89 dBQ and noise 16 pA/√Hz. The powerconsumption of each TIA 804 can be approximately 0.551 mW.

The single to differential converter 806 can convert the single endedsignal of the TIA 804 to a fully differential signal. Fully differentialsignal processing reduces the effect of external noises, clock injectionand even order harmonic and increases dynamic range. FIG. 12 illustratesthe STD circuit schematic. FIG. 13 illustrates an example of the antialiasing LPF 808. In this example, a fully differential 4th order biquadtunable gm-C low pass filter is used to limit the filter is designed toband limit the TIA's 804 output signal. Each LPF 808 can have a −3 dBbandwidth of 11 MHz and consume 6.7 mW power. FIG. 13 shows the biquadstructure of the gm-C filter. Other examples can eliminate the LPF 308by adjusting the CMUT array element frequency response.

The passive mixer 810 can be designed to modulate the outputs ofmultiple (in this example 8) different LPFs 808 with orthogonal carrierfrequency of 40 MHz, 80 MHz, 120 MHz and 160 MHz. FIG. 14 shows anexample of an I/Q passive mixer 810 structure. The BPF 812 can beutilized to eliminate the harmonic generated by the passive mixer 810.In an example, four 8th order tunable gm-C biquad band-pass filters 812with center frequency 40 MHz, 80 MHz, 120 MHz and 160 MHz can be used.The bandwidth of each BPF 812 can be ˜30 MHz. The power consumption ofeach BPF 812 can be 23 mW. The outputs of the BPFs 812 can be combinedusing a power spectrum combiner circuit known in the art. Aninstrumentation amplifier can also be used to convert the differentialoutput of combiner circuit to single ended.

The buffer 814 can be a current feedback source degenerated push-pulltype. A simulated bandwidth of the buffer 814 can be ˜350 MHz for a loadof 35pfll1MΩ. Further, a clock generator 816 can be provided. In anexample, two clocks of 240 MHz and 320 MHz are supplied externally.Using divide by 3, and a walk-in-ring oscillator circuit, the 40 MHzorthogonal carrier signals are generated from 240MHz signals. Usingdivide by 2, and a walk-in-ring oscillator circuit, the 80 MHzorthogonal carrier signals can be generated from 320 MHz. 120 MHz and160 MHz orthogonal carriers can be generated by feeding 240 MHz and 320MHz clocks directly to walk-in-ring oscillator respectively.

A simulation of the OFDM multiplexing component 800 was performed in aCADENCE environment using Specter circuit simulation tool. Post layoutsimulation was performed of the individual blocks and the entire system.FIG. 15 shows the simulation result for the TIA 804 showing 89 dBQ andbandwidth of 20 MHz, as designed for one example. To simulate thefunctionality to single to differential converter a 7 MHz centerfrequency 80% bandwidth Gaussian pulse was applied to the input which isshown in FIG. 16A. The outputs of the differential converter are shownin FIG. 16B, indicating close to the desired 180 degree phase shift.FIG. 17 illustrates the frequency response of the designed LPF 808 whichhas almost −80 dB/decade sharp roll off and 11 MHz bandwidth and FIG. 18illustrates the frequency response of the designed BPFs 812. FIG. 18illustrates that the center frequencies of the filters are 40 MHz, 80MHz, 120 MHz and 160 MHz and bandwidth of each filter is ˜30 MHz.

FIGS. 19A-19C illustrate some of the performance validation results ofone example. To validate the performance of the analog OFDM anddemodulation scheme the chip was first simulated with two 7 MHz 80%bandwidth Gaussian pulses shown FIG. 19A representing inputs from twochannels. The two pulses were up-converted with orthogonal 40 MHzcarriers shown FIG. 19B. The up-converted signals were de-modulatedwhich are shown in FIG. 19C. It is clear from FIG. 19C that relativeamplitude and phase of the signals were preserved.

Although not detrimental, cross talk between the channels, lower than−20 dB, is also observed which may be due to the non-ideal single todifferential, differential to single conversion or mixing. To verify thefunctionality of the entire designed chip 7 MHz 80% BW Gaussian pulseswith different phase were applied to all 8 channels and modulated. FIG.20A shows spectrum of a single channel input spectrum. FIG. 20B showsspectrum of two Gaussian pulses which are up-converted to 40 MHz. Thespectrum of all the up-converted signals is shown in FIG. 20C. Thespectra of the modulated signals are not identical to each other due tothe different initial phases. During this simulation the total powerconsumed by the chip was ˜160 mW.

While the present disclosure has been described in connection with aplurality of exemplary aspects, as illustrated in the various figuresand discussed above, it is understood that other similar aspects can beused or modifications and additions can be made to the described aspectsfor performing the same function of the present disclosure withoutdeviating therefrom. For example, in various aspects of the disclosure,methods and compositions were described according to aspects of thepresently disclosed subject matter. However, other equivalent methods orcomposition to these described aspects are also contemplated by theteachings herein. Therefore, the present disclosure should not belimited to any single aspect, but rather construed in breadth and scopein accordance with the appended claims.

1. A CMUT on CMOS chip for imaging applications, comprising: a pluralityof CMUT transmit (“Tx”) elements transmitting an imaging pulse; aplurality of CMUT receive (“Rx”) elements, disposed in proximity to theplurality of Tx elements, receiving the imaging pulse, and each Rxelement generating a receiver signal; and a plurality of electronicsinterfacing with the Tx elements and the Rx elements, comprising: amultiplexing component receiving at least half of the generated receiversignals simultaneously and reducing a number of output signals, based onthe receiver signals, by a ratio between 15-to-1 and 50-to-1; and abeamforming component communicating with Tx elements.
 2. The CMUT onCMOS chip of claim 1, wherein the multiplexing component uses frequencydivision multiplexing (“FDM”).
 3. The CMUT on CMOS chip of claim 2,wherein the multiplexing component comprises: a feedback transimpedanceamplifier (“TIA”) receiving the receiver signal from the Rx elements; amixer receiving a first signal from the TIA; a band pass filter (“BPF”)receiving a second signal from the mixer; an adder receiving a thirdsignal from the BPF; and a buffer receiving a fourth signal from theadder and producing the output signal.
 4. The CMUT on CMOS chip of claim1, wherein the beamforming component comprises: a high voltage pulserelectrically connected to the Tx elements; and a timing and codingcircuit electrically connected to the high voltage pulser.
 5. The CMUTon CMOS chip of claim 1, wherein the multiplexing component produces theoutput signal using time division multiplexing (“TDM”).
 6. The CMUT onCMOS chip of claim 5, wherein the multiplexing component comprises: afeedback transimpedance amplifier (“TIA”) receiving the receiver signalfrom the Rx elements; a time-gain compensation (“TGC”) circuit receivinga first signal from the TIA; a time division multiplexing switchreceiving a second signal from the TGC; and a buffer receiving a thirdsignal from the time division multiplexing switch and producing theoutput signal.
 7. The CMUT on CMOS chip of claim 5, wherein the timedivision multiplexing switch samples the second signal at a rate greaterthan the Nyquist rate.
 8. The CMUT on CMOS chip of claim 5, wherein theplurality of electronics comprises: a de-multiplexer receiving the RFoutput signal; and a low pass filter receiving a first signal from thede-multiplexer and outputting a non-combined output signal.
 9. The CMUTon CMOS chip of claim 1, wherein the field of view is between 90°×90° atapproximately 5 cm depth and 45°'45° at approximately 15 cm depth. 10.The CMUT on CMOS chip of claim 1, wherein the multiplexing componentreduces the number of output signals, based on the receiver signals, bya ratio of at least 16-to-1.
 11. The CMUT on CMOS chip of claim 1,wherein the multiplexing component uses orthogonal frequency divisionmultiplexing (“OFDM”).
 12. The CMUT on CMOS chip of claim 11, whereinthe multiplexing component comprises: a feedback transimpedanceamplifier (“TIA”) receiving the receiver signal from the Rx elements; asingle to differential converter (“STD”) receiving a first signal fromthe TIA; a low band pass filter (“LPF”) receiving a second signal fromthe STD; a mixer receiving a third signal from the LPF; a band passfilter (“BPF”) receiving a fourth signal from the mixer; and a bufferreceiving a fifth signal from the adder and producing the output signal.13. The CMUT on CMOS chip of claim 1, wherein at least one of theplurality of Tx elements; the plurality of Rx elements; and theplurality of electronics is on a separate CMUT on CMOS chip in a stackedconfiguration.
 14. A system for intracardiac imaging, comprising: anintracardiac echography catheter having an internal volume, a proximalend and a distal end, comprising: an atraumatic tip disposed on thedistal end of the catheter; a CMUT on CMOS volumetric imaging chipdisposed proximal of the tip, comprising: a plurality of CMUT transmit(“Tx”) elements transmitting an imaging pulse; a plurality of CMUTreceive (“Rx”) elements, disposed in proximity to the plurality of Txelements, receiving the imaging pulse, and each Rx element generating areceiver signal; and a plurality of electronics interfacing with the Txelements and the Rx elements, comprising: a multiplexing componentreceiving at least half of the generated receiver signals simultaneouslyand reducing a number of output signals, based on the receiver signals,by a ratio between 15-to-1 and 50-to-1; and a beamforming componentcommunicating with Tx elements; and a cable lumen disposed within thevolume and configured to receive a plurality of output signal cablesfrom the chip.
 15. The system of claim 14, wherein a size of thecatheter is approximately 6 French to approximately 10 French.
 16. Thesystem of claim 14, wherein the catheter further comprises a coolinglumen cooling the catheter from RF induced heating when the catheter isused in conjunction with an MRI scan.
 17. The system of claim 14,wherein the catheter further comprises a pair of inductively coupledcoils proximal the atraumatic tip wherein one coil is proximal of thechip and the second coil is distal of the chip, wherein the coils areconfigured to be visible during MRI operation.
 18. The system of claim14, wherein the catheter further comprises: a body comprising at leastone MRI safe material; at least one direction wire disposed in adirectional wire lumen, the wire comprising at least one MRI safematerial; wherein the catheter is configured to be used during at leastone of interlaced and simultaneous MRI and ultrasound operation.
 19. Thesystem of claim 18, wherein the at least one MRI safe materials are alsox-ray safe materials.
 20. The system of claim 14, wherein themultiplexing component produces the output signal using at least one offrequency division multiplexing (“FDM”), time division multiplexing(“TDM”), and orthogonal frequency division multiplexing (“OFDM”).